Swellable particles for drug delivery

ABSTRACT

Swellable particles for delivering a working agent to the pulmonary system comprise a plurality of biodegradable particles each formed from a polymer network, each of the plurality of biodegradable particles having a mass mean aerodynamic diameter not exceeding 5 μm, the particles being swellable by hydration to a size that is greater than 6 μm volume mean diameter, and a working agent entrapped in the polymer network of each of the plurality of biodegradable particles.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.13/601,532, filed Aug. 31, 2012, which is a divisional of U.S.application Ser. No. 11/732,489, filed Apr. 3, 2007, which is anon-provisional of and claims the benefit of priority under 35 U.S.C.§119(e) to U.S. Provisional Application No. 60/788,942, filed Apr. 4,2006, all of which are incorporated herein by reference in theirentirety.

BACKGROUND OF THE INVENTION

The present invention relates generally to swellable biodegradableparticles for use in delivering a therapeutic or other agent to thepulmonary system and, more particularly, to swellable biodegradableparticles that have dehydrated sizes for delivery to the pulmonarysystem and that swell on hydration in the pulmonary tract to largersizes to achieve controlled release of a drug or other agent from theparticle structure.

BACKGROUND OF THE INVENTION

Delivery of therapeutic agents to the pulmonary system has been used forthe treatment of local lung diseases such as asthma, cystic fibrosis,and chronic obstructive pulmonary disease (A. J. Hickey, editor,Inhalation Aerosols: Physical and Biological Basis for Therapy, NewYork: Marcel Dekker, Inc. 1996). Relative to systemic oral or injectiondrug delivery, local delivery of respiratory drugs to the lungs providesadvantages because it: (1) requires smaller doses of the drug; and (2)minimizes systemic toxicity by allowing delivery directly to the site ofthe disease. Delivery of systemically acting agents has also beeninvestigated, such as for the administration of proteins and peptides(e.g. insulin) as described by Patton et al., in “Inhaled insulin”, Adv.Drug. Deliv. Rev. 35, pp. 235-247 (1999). However, pulmonary delivery ofdrugs is limited by several major issues including poor efficiency ofdeposition in the respiratory tract and excessive removal of the drug bythe oropharyngeal cavity, poor control over the site of deposition ofthe drug within the respiratory tract, poor reproducibility of dosingdue to the dependence on breathing patterns of the patient, and toorapid clearance and/or absorption of the drug from the pulmonary systempotentially resulting in inappropriate drug concentrations at the targetsite and even toxic effects.

A controlled release delivery system for drugs delivered locally to thelung would provide a very desirable method to effectively treatrespiratory and systemic diseases. Moreover, controlled release ofrespiratory drugs may offer significant clinical benefit to millions ofpatients with respiratory disease by allowing them to take treatmentsfor such diseases as asthma less frequently and to receive moreprolonged and controlled relief. Controlled delivery of drugs to thelung also offers the potential for improved safety by moderating thedrug peaks and troughs of immediate release drugs, which can cause addedtoxicity or reduced efficacy. Also, controlling the release of two ormore therapeutic agents from a single particle system delivered to thepulmonary system would have significant benefits for co-localization ofthe agents within the respiratory tract. The likelihood of synergism oradditive effects between agents would be significantly increased.

Currently available pulmonary delivery systems are not ideal, deliveringinaccurate doses, requiring frequent dosing and losing significantamounts of drug in the delivery process. Most asthma drugs delivered viainhalation are immediate-release formulations that must be inhaledmultiple times per day (Cochrane et al. Inhaled Corticosteroids forAsthma Therapy: Patient Compliance, Devices, and Inhalation TechniqueChest. 117, pp. 542-550, 2000). This frequent inhalation tends todiscourage patient compliance. When patients forget to take theirmedicine they may experience complications which may result in increasedemergency room visits and hospitalizations. In a recent analysis ofpublished studies of patient compliance with asthma medications,patients took the recommended doses of medication on only 20 to 73% ofdays (Cochrane et al. Inhaled Corticosteroids for Asthma Therapy:Patient Compliance, Devices, and Inhalation Technique, Chest. 117, pp.542-550, 2000). The percentage of under-use days ranged from 24 to 69%.In addition, immediate release formulations often deliver drug levelsthat peak and trough, causing undesirable toxicity or inadequateefficacy.

Although promising, inhaled formulations face difficult challenges inmaintaining effective drug concentration in the lungs for extendedperiods. Factors contributing to the short duration of drug actionfollowing pulmonary delivery include: (1) the rapid mucociliaryclearance rates (approximately 1.7-4.9 mm/min) resulting in a very shorthalf-life for inhaled particles (approximately 0.5-2 hr) (Lansley, A.B., 1993. Mucociliary clearance and drug delivery via the respiratorytract. Adv Drug Del Rev. 11, 299-327); (2) phagocytosis of particles bythe alveolar macrophages; and; (3) rapid absorption of drug molecules(Mw<1000 Da) to the systemic circulation with a mean half-time forabsorption of <2 hr.

The pulmonary region has several particle clearance mechanisms. Therelative importance of each clearance mechanism varies depending on thephysicochemical properties of the particle. Particle retention in thepulmonary region is longer than that of the ciliated airways. Afterdeposition, uptake of particles by alveolar macrophages is very rapid.An initial fast phase of clearance is related to phagocytosis byalveolar macrophages.

There are limited technologies available to circumvent the naturalclearance mechanisms of the airways that largely prevent sustainedrelease particles from being effective. A number of prior artreferences, including, but not limited to U.S. Pat. No. 6,136,295 toEdward, et al.; and U.S. Pat. No. 6,730,322 to Bernstein, et al.,describe particles that have been designed to have low densities (largeporous particles). Although geometrically large, those particles areaerodynamically much smaller.

Generally to achieve sustained release, particles must be delivered tothe airways and avoid mucociliary clearance, uptake by alveolarmacrophages, and prevention of rapid absorption from the lung. Avoidingmucociliary clearance can be achieved by avoiding particle deposition inthe tracheobronchial region where ciliated epithelia are present.Generally an aerodynamic particle size must be less than around 5 μm toaccomplish this. Once particles are deposited in the peripheral airwayswhere the mucociliary clearance mechanism is not present, particles mustavoid alveolar macrophage uptake that can rapidly clear therapeuticcompounds. Avoidance of macrophages can be accomplished by (1) creatingparticles that are not recognizable as foreign particulates (stealthparticles); (2) providing particles that are physically too large to beengulfed by macrophages or which delay engulfment; or (3) providingparticles that are too small to be recognized by macrophages(nanoparticles).

Current sustained release pulmonary systems as described by the citedprior art generally comprise large porous particle technologies. Themain problems with these systems is the low drug loading possible in theparticle matrix, the special physicochemical properties of the drugrequired for inclusion in these particle systems, and the limits on howlong drug may be sustained. The present invention overcomes theseproblems by using swelling particles to improve sustained release. Theswelling particles of the present invention include the drug or otherworking agent being delivered on and/or in a biocompatible andbiodegradable swellable matrix that preferably enables deep lungdelivery and avoids clearance by the alveolar macrophages. In addition,the matrix materials can be modified to modulate the drug releasecharacteristics or to improve compatibility of the drug with the matrixsystem.

SUMMARY OF THE INVENTION

The present invention provides improved swellable particles for deliveryto the pulmonary system, and to a method for their incorporation andadministration of a working agent, such as including but not limited toa therapeutic agent, diagnostic agent, prophylactic agent or imagingagent. The swellable particles include dehydrated (dry) aerodynamicparticle diameters to enable delivery to the respiratory tract, such asfor example to the tracheo-bronchial airways of the upper respiratorytract and/or to the alveolar regions of the deep lung, and hydratedparticle diameters that are greater than 6 μm volume mean diameter toretard or prevent their phagocytosis by the macrophages present inairways of the respiratory tract.

In an illustrative embodiment of the invention, the dehydrated (dry)particles are made of a biodegradable material, have a mass medianaerodynamic particle diameter between 0.5 μm and 5 μm, and are capableof swelling to a hydrated geometric particle diameter greater than 6 μmvolume mean diameter. In a preferred embodiment of the invention, atleast 90%, more preferably 95% to 99%, of the particles have anaerodynamic particle diameter not exceeding 5 μm and swell to a size ofgreater than 6 μm volume mean diameter. The particles may be formed ofbiodegradable materials such as including, but not limited to, abiodegradable natural or synthetic polymer, a protein, a carbohydrate,or combinations thereof. For example, the particles may be formed of amulti-branched polyethylene glycol (PEG) hydrogel polymer. Otherexamples include particles formed of biodegradable polymers such asdextrans, hydroxyethylmethylacrylate, or other biocompatible andbiodegradable swellable polymeric systems. The swellable particles canbe used for enhanced delivery of one or more working agents to theairways of the respiratory tract, including to the alveolar region ofthe lung. The particles incorporating one or more working agents may beeffectively aerosolized for administration to the respiratory tract topermit systemic or local delivery of a wide variety of therapeutic andother agents. They optionally may be co-delivered with larger carrierparticles (not carrying a therapeutic or other agent) which have forexample a mean diameter ranging between about 50 μm and 150 μm.

The present invention is advantageous in that the dehydrated (dry)particles possess an aerodynamic diameter such that a) they are able toreach one or more target regions of the respiratory tract, including thetracheo-bronchial airways of the upper respiratory tract and/or to thealveolar regions of the deep lung, b) they can deliver a payload of oneor more working agents without premature release, c) they are swellableby hydration in the airways to a size that retards or prevents theiruptake and removal by macrophages, and d) they can provide controlledrelease of the working agent(s) at predictable rates followinghydration.

Other features and advantages of the present invention will becomeapparent from the following detailed description.

DETAILED DESCRIPTION OF THE INVENTION

The present invention provides swellable, biodegradable particles forimproved delivery of therapeutic and other working agents to therespiratory tract. The working agents which can be delivered via theparticles include, but are not limited to, a therapeutic agent,diagnostic agent, prophylactic agent, imaging agent, or combinationsthereof.

In an illustrative embodiment of the invention, the swellable particlesinitially comprise dehydrated (dry) powder particles having mass medianaerodynamic particle diameter of 5 μm or less to enable delivery to therespiratory tract, such as for example to the tracheo-bronchial airwaysof the upper respiratory tract and/or to the alveolic regions of thedeep lung, and having hydrated particle diameter that is greater than 6μm volume mean diameter to retard or prevent their phagocytosis by themacrophages present in airways of the respiratory tract. The dehydrated(dry) particles typically have a mass median aerodynamic diameterbetween 0.5 μm and 5 μm, and typically are capable of swelling to ahydrated geometric diameter greater than 6 μm to 50 μm volume meandiameter. At least 90%, more preferably 90% to 95%, of the particleshave an aerodynamic diameter of 5 μm or less and swell to a size ofgreater than 6 μm volume mean diameter.

The mass median aerodynamic diameter (MMAD) is typically obtained fromconventional aerosol sizing instruments, such as cascade impactorsand/or time of flight instruments such as the TSI Aerodynamic ParticleSizer (TSI Incorporated, Shoreview, Minn.). This size determinationoccurs where 50% of the mass of the particles, when classified by theiraerodynamic size, are below this diameter (i.e. the MMAD). That is, 50%of the mass of particles have a diameter lower than the MMAD and 50% ofthe mass of particles have a diameter higher than the MMAD. This measureof particle size converts the particle in question (which can havedifferent densities, shapes, surface and aerodynamic drag) into a spherehaving a density equal to 1 and provides an equivalent sphere diameter,even though the particle may have a flake, acicular or othernon-spherical shape. For example, the aerodynamic diameter is defined asthe diameter of the spherical particle with a density of 1 g/cm³ (thedensity of a water droplet) that has the same settling velocity as theparticle and is given by the following equation (see Hinds, W. I.,“Uniform Particle Motion,” in Aerosol Technology 1999, pp. 42-74, JohnWiley and Sons Inc.): d_(a)=d_(p)(ρ_(p))^(1/2) where d_(p) is thediameter of the particle and ρ_(p) is its density in g/cm³. Theaerodynamic diameter may be thought of as the diameter of a sphericaldroplet of water possessing the same aerodynamic properties as theparticle. For example, if a particle has an aerodynamic diameter of 1μm, it acts aerodynamically identical to a 1 μm water droplet regardlessof the particle's actual size, shape, or density. By adjusting the meandiameter and density of an aerosol population, the particles can betailored to possess the exact aerodynamic diameter necessary fordelivery to a specified lung region. The diameters of the swellableparticles in a sample can range depending upon on factors such asparticle composition and methods of synthesis. The distribution of sizeof particles in a sample can be selected to permit optimal depositionwithin targeted sites within the respiratory tract.

The equivalent sphere diameter allows one to directly compare particleswith different particle geometries and compare them only on the basis ofaerodynamics, which is the functional size characteristic of importancefor delivery of the particles to the respiratory tract. The MMAD isselected to describe the dehydrated (dry) particles because it is usefulin predicting the deposition of dry powder particles within the airways.

The volume mean diameter is used to describe the hydrated particlesbecause the particles will have swollen in size after hydration in theairways. The volume mean diameter will increase from the dry state tothe hydrated state regardless of the particle shape. The volume meandiameter is also an equivalent sphere diameter whereby the particle inquestion is converted to a sphere of equivalent volume, even though theparticle may have a flake, acicular or other non-spherical shape. Interms of functionality, the volume of the particle is important becauseit is related to how well the macrophage cells in the airways can clearthe particles; i.e. their volume is important for retarding or avoidingthe clearance mechanisms of the respiratory tract. The volume meandiameter can be determined by testing as follows: using hydrogelparticles dispersed in buffer solution, a laser light scatteringinstrument can be used to measure volume diameter changes in theparticle geometry upon hydration. A low energy liquid dispersionattachment may also be used to minimize particle aggregation.Alternatively, particle swelling can be quantified via confocalmicroscopy.

The swellable particles may be formed from any biodegradeable, andpreferably biocompatible polymer, copolymer, or blend, which is capableof forming particles with a mass median aerodynamic diameter between 0.5and 5 μm, but can also swell to a geometric diameter of greater than 6μm volume mean diameter. For purposes of illustration and notlimitation, the particles can be formed of a swellable biodegradablenatural polymer, synthetic polymer, protein, carbohydrate, orcombinations thereof. For example, the particles may be formed of amulti-branched polyethylene glycol (PEG) hydrogel polymer. Otherexamples include particles formed of biodegradable polymers such asdextrans, hydroxyethylmethylacrylate, or other biocompatible andbiodegradable swellable polymeric systems.

For purposes of further illustration and not limitation, the swellableparticles also can be made from bulk eroding hydrogel polymers, such asthose based on polyesters including poly(hydroxy acids) can be used inpractice of the invention. Moreover, surface eroding polymers such aspolyanhydrides may be used to form the swelling particles. For example,polyanhydrides such as poly[(p-carboxyphenoxy)-hexane anhydride] (PCPH)may be used. For example, polyglycolic acid (PGA) or polylactic acid(PLA) or copolymers thereof may be used to form the swellable particles,wherein the polyester has incorporated therein a charged orfunctionalizable group such as an amino acid as described below. Otherpolymers include polyamides, polycarbonates, polyalkylenes such aspolyethylene, polypropylene, poly(ethylene oxide), poly(ethyleneterephthalate), poly vinyl compounds such as polyvinyl alcohols,polyvinyl ethers, and polyvinyl esters, polymers of acrylic andmethacrylic acids, celluloses and other polysaccharides, and peptides orproteins, or copolymers or blends thereof which are capable of formingswellable particles described above. Polymers may be selected with or bemodified to have the appropriate stability and degradation rates in vivofor different controlled drug delivery applications.

Features of the swellable particle which can contribute to swellinginclude degree of polymer cross-linking, monomer size, and porosity. Arough particle surface texture also can reduce particle agglomerationand provide a highly flowable powder, which is ideal for aerosolizationvia dry powder inhaler devices, leading to lower deposition in themouth, throat and inhaler device. Moreover, administration of theswellable particles to the lung by aerosolization permits deep lungdelivery of therapeutic aerosols where the particles can swell afterhydration on the airway surfaces). In order to serve as efficient drugcarriers in drug delivery systems, the swellable particles preferablyare biodegradable and biocompatible, and optionally are capable ofbiodegrading at a controlled rate for delivery of a drug.

In an illustrative embodiment of the invention, biodegradeable and/orbiocompatible hydrogel particles are formed from acrylated 8-arm PEG (20kDa) crosslinked with dithiothreitol, as described by Hubbel et al.Journal of Controlled Release 76:11-25 (2001), the teachings of whichare incorporated herein by reference. Alternate methods of crosslinkingthe acrylated polymers include photopolymerization, other covalentcrosslinking methods (polycysteine), ionic crosslinking, and physicalcrosslinking (entanglements between highly branched and high molecularweight polymers).

In the synthesis, both the molecular weight, degree of branching (e.g.8-arm to 4-arm), and the concentrations of the reactants may be alteredto change the pore size of the hydrogels, and thereby adjust the releaserate of the therapeutic within the polymer matrix.

Alternatively, the biocompatible hydrogel particles may be formed ofmultiple polymer molecules, copolymeric hydrogels, chosen to grantspecific and advantageous characteristics to the system. In onepreferred embodiment, hydrogels are constructed from a block copolymerconfiguration of repeating units of poly-lactic acid (PLA) andpolyethylene glycol. The PLA confers rapidly hydrolyzable ester bonds,whereas the PEG backbone prevents both the rapid degradation of thepolymer and adsorption of proteins to the hydrogel surface andsubsequent removal by the immune system.

The swellable particles can be made using a variety of particle-formingprocesses and in a variety of particle shapes. For example, swellingpolymeric particles can be prepared using spray drying, solventevaporation, polymer micronization, and other methods well known tothose of ordinary skill in the art. Swellable particles comprisinghydrogels can be synthesized during emulsification with an aqueous phasewith a non-aqueous phase to form microspheres of hydrogel that can bemodulated in size by changing parameters of the emulsion (e.g.non-aqueous phase composition, concentration of reactants, mixing speedand shear introduced into the emulsion, presence of surfactants andsurfactant types, etc). Swellable particles comprising hydrogels alsocan be synthesized during spraying so that the hydrogel particles formand are cross linked while dispersed as a droplet. Alternately, disks,spheres, cubes, irregular shapes, thin films of hydrogels can besynthesized and then broken into (comminuted) small respirable,swellable particles using milling and micronization methods. Thecomminuted particles can have a flake, acicular or other non-sphericalshape.

For example, the dried hydrogels made can be broken down intomacroparticles using a rotary blade mill (M 20 Universal mill, IKA®Werke GmbH, Germany). Size reduction to narrowly dispersed micron sizeparticles suitable for inhalation can then be performed using a fluidenergy mill that uses impinging air jets to finely micronize thematerial. Particle size of the resultant material is controlled by theparameters of fluid energy milling (air pressures) and can also beseparated using an air classifier. Particle size can be determined usinga Sympatec Helos laser diffraction instrument and also conventionalcascade impaction techniques. Milled particles exhibit much faster ratesof swelling than unmilled particles as a result of the increased surfacearea available for water uptake.

The swellable particles may be fabricated or separated, for example bysieving or air separation methods, to provide a particle sample with apreselected size distribution to provide the desired MMAD in the drypowder state. As mentioned above, the selected range within which acertain percentage of the particles must fall preferably is controlledsuch that at least 90%, or even more preferably 95% or 99%, have anaerodynamic particle diameter between 0.5 μm and 5 μm.

A particular process for making swellable particle starts with hydratedfilms or hydrogel masses, which can then be processed for particle sizereduction and micronization. For example, the hydrated films or hydrogelmasses can be extruded through a fine orifice or mesh to reduce particlesize. The particles can then be dehydrated. Alternatively, particles canbe produced after dehydration of the hydrogel films. Dehydration isachieved using methods such as a vacuum oven drying, lyophilization,solvent displacement by volatiles, among others. The dried hydrogels canbe broken down into particles using a rotary blade mill (M 20 Universalmill, IKA® Werke GmbH, Germany). Size reduction to narrowly dispersedmicron size particles suitable for inhalation can then be performedusing a fluid energy mill that uses impinging air jets to finelymicronize the material.

The swellable particles can be used for enhanced delivery of one or moreworking agents to the airways of the respiratory tract, including thealveolar region of the lung. The particles incorporating one or multipleworking agents may be effectively aerosolized for administration to therespiratory tract to permit systemic or local delivery of a wide varietyof therapeutic agents. For example, particle size of an inhaled aerosolis the primary determinant of the deposition pattern within the airways.They optionally may be co-delivered with larger carrier particles (i.e.not carrying a working agent) which have for example a mean diameterranging between about 50 μm and 150 μm.

Incorporation of a therapeutic or other working agent within theparticle can be accomplished using a variety of methods. For example,inclusion/incorporation of a working agent inside the particlescomprising hydrogels described herein can be achieved by entrapping theworking agent in the hydrogel polymer network so as to control therelease of the drug (or other working agent) from the particle atspecifically desired rates. Entrapment can be achieved by performing thecross linking of the polymers in the presence of the working agent (e.g.drug) such that the working agent (e.g. drug) is entrapped within thepolymer network that forms the hydrogel network. Entrapment also can beachieved by performing the cross linking of the polymers prior toplacing the hydrogels in the presence of the working agent (e.g. drug)such that the working agent is entrapped within the polymer network bydiffusing into the hydrogel. Moreover, multiple working agents (e.g.drugs) can be loaded into the same swellable particles using the same ordifferent methods since beneficial release rates may be achieved byloading differently or using the same methods.

The therapeutic agent (or other working agent) to be loaded into theswellable hydrogel particles can take various forms including a drug insolution such that the drug is a molecular dispersion throughout thehydrogel particle, a drug in suspension, a colloidal dispersion, ananoparticle system dispersed throughout the hydrogel particle, or adrug in liposomes, lipid dispersions, nanocapsules, polymericnanoparticles, etc dispersed throughout the hydrogel particle.

For purposes of further illustration and not limitation, incorporationof a therapeutic agent (or other working agent) within the particle canbe accomplished by the following methods:

a) Encapsulation of the therapeutic agent within a nanoparticle andplacement of the nanoparticle within the particle. For example,Ibuprofen can be encapsulated within Lecithin (phophotidylcholine)nanoparticles 200-400 nm in diameter. These nanoparticles were preparedusing the method of Chen et al, 2002 (Chen, X., Young, T. J., Sarkari,M., Williams, R. O., Johnston, K. P., Preparation of cyclosporine Ananoparticles by evaporative precipitation into aqueous solution,International Journal of Pharmaceutics 242, (2002) 3-14). Thenanoparticles are incorporated into the swellable particles by physicalentrapment in the hydrogel network by performing crosslinking reactionsin the presence of the nanoparticles in the reactant solution of the 20kDa 8-arm acrylated PEG with dithiothreitol as described below for theillustrative embodiment. Nanoparticles can be designed to havedifferential rates of release from the swellable particles based ontheir relative sizes, hydrophilic nature, ionic properties, anddiffusion coefficients.

b) Direct encapsulation of the molecule within the matrix of the polymernetwork (e.g. PEG hydrogel network) of the swellable particle. Forexample, Rhodamine therapeutic agent can be trapped within the polymernetwork (PEG hydrogel network) during the crosslinking of the 20 kDa8-arm acrylated PEG with dithiothreitol as described below for theillustrative embodiment.

c) Attachment of the therapeutic agent to the polymer network itself(PEG hydrogel network) through chemical interactions (covalent, ionic,and hydrogen bonds). For example, N-acetylcysteine mucolytic agent canbe attached to the polymer network of 8-arm acrylated PEG crosslinkedwith dithiothreitol through covalent bonds between the thiol group onthe N-acetylcysteine and one thiol group on the dithiothreitol with theother thiol group of dithiothreitol bonded to an acrylate group of thepolymer network. The N-acetylcysteine mucolytic agent can be reversiblycovalently bound to the hydrogel network of the particle such that theN-acetylcysteine functions to decrease the viscosity of the mucus bydisrupting the disulfide bonds formed between adjacent cysteineresidues. This disruption of mucus disulfide bonds is readily achievedsince the disulfide bonds are transferred between cysteine residues.This facilitates prolonged and localized mucolytic release around thehydrogel particle structure, increasing transport rates through the CFmucus environment.

Using these encapsulation methods, numerous therapeutic and otherworking agents, ranging from small and hydrophilic to large andhydrophobic, may be incorporated into the swellable particles for theaerosolized treatment of cystic fibrosis, lung cancer, asthma, chronicobstructive pulmonary disease (COPD), acute bronchitis, emphysema,tuberculosis, or systemic diseases. For example, loading of Rhodaminetherapeutic agent in PEG hydrogel particles described herein has beenachieved during polymerization or after the particles were made. Highloading of this drug was obtained using both methods. For example,approximately 35% w/w Rhodamine therapeutic agent was present afterwashing surface drug from the hydrogel particles.

Any of a variety of therapeutic treating agents, prophylactic agents,diagnostic agents, imaging agents such as radio-isotopes, or otheractive working agents can be incorporated within the swellableparticles. The swellable particles can be used to locally orsystemically deliver a variety of therapeutic agents to the respiratorytract. Examples of working agents include synthetic inorganic andorganic compounds, proteins and peptides, polysaccharides and othersugars, lipids, and nucleic acid sequences having therapeutic,prophylactic, diagnostic or imaging activities. Nucleic acid sequencesinclude genes, antisense molecules which bind to complementary DNA toinhibit transcription, and ribozymes. The working agents to beincorporated can have a variety of biological activities, such asvasoactive agents, neuroactive agents, hormones, anticoagulants,immunomodulating agents, cytotoxic agents, prophylactic agents,antibiotics, antivirals, antisense, antigens, and antibodies. In someinstances, the proteins may be antibodies or antigens which otherwisewould have to be administered by injection to elicit an appropriateresponse. Compounds with a wide range of molecular weight can beencapsulated, for example, between 100 and 500,000 μm per mole.

Proteins are defined as comprising 100 amino acid residues or more;peptides are less than 100 amino acid residues. Unless otherwise stated,the term protein refers to both proteins and peptides. Examples includeinsulin and other hormones. Polysaccharides, such as heparin, can alsobe administered.

The swellable polymeric aerosols are useful as carriers for a variety ofinhalation therapies. They can be used to encapsulate small and largedrugs, release encapsulated drugs over time periods ranging from hoursto months, and withstand extreme conditions during aerosolization orfollowing deposition in the lungs that might otherwise harm theencapsulated therapeutic.

For example, the swellable particles may include a therapeutic agent forlocal delivery within the lung, such as agents for the treatment ofasthma, emphysema, or cystic fibrosis, or for systemic treatment. Forexample, genes for the treatment of diseases such as cystic fibrosis canbe administered, as can beta agonists for asthma. Other specifictherapeutic agents include, but are not limited to, insulin, calcitonin,leuprolide (or LHRH), G-CSF, parathyroid hormone-related peptide,somatostatin, testosterone, progesterone, estradiol, nicotine, fentanyl,norethisterone, clonidine, scopolamine, salicylate, cromolyn sodium,salmeterol, formeterol, albuterol, and vallium.

The particles including a therapeutic agent may be administered alone orin any appropriate pharmaceutical carrier, such as an inert sugarparticle system typically used in a powder inhaler, for administrationto the respiratory system. They can be co-delivered with larger carrierparticles (not including a therapeutic agent) possessing mass meandiameters for example in the range 50 μm to 150 μm.

Aerosol dosage, formulations and delivery systems may be selected for aparticular therapeutic application, as described, for example, in Gonda,I. “Aerosols for delivery of therapeutic and diagnostic agents to therespiratory tract,” in Critical Reviews in Therapeutic Drug CarrierSystems, 6:273-313, 1990; and in Moren, “Aerosol dosage, forms andformulations,” in: Aerosols in Medicine. Principles, Diagnosis andTherapy, Moren, et al., Eds, Esevier, Amsterdam, 1985, the disclosuresof which are incorporated herein by reference.

The relatively large size of swollen aerosol particles deposited in thedeep lungs minimizes potential drug losses caused by particlephagocytosis. The swellable polymeric matrix also facilitates as atherapeutic carrier to provide the benefits of biodegradable polymersfor controlled release in the lungs and long-time local action orsystemic bioavailability. Denaturation of macromolecular drugs can beminimized during aerosolization since macromolecules are contained andprotected within a polymeric matrix shell. Coencapsulation of peptideswith peptidase-inhibitors can minimize peptide enzymatic degradation.

For purposes of still further illustration and not limitation, theswellable particles can include a working agent that comprises amucolytic agent alone or together with another working agent such asantibiotic agent, a cytotoxic agent, other mucolytic agents, an RNAinterfering agent which includes siRNA and miRNA, a gene, orcombinations thereof. The working agent also can comprise multiplecytotoxic agents, a cytotoxic agent and an RNA interfering agent, orother combinations of working agents for purposes of furtherillustration.

In comparison to non-swellable particles, the swellable particlespursuant to the present invention also can potentially more successfullyavoid phagocytic engulfment by alveolar macrophages and clearance fromthe lungs, due to size exclusion of the particles from the phagocytes'cytosolic space. Phagocytosis of particles by alveolar macrophagesdiminishes precipitously as particle diameter increases beyond 3 μmKawaguchi, H. et al., Biomaterials 7: 61-66 (1986); Krenis, L. J. andStrauss, B., Proc. Soc. Exp. Med., 107:748-750 (1961); and Rudt, S, andMuller, R. H., J. Contr. Rel., 22: 263-272 (1992). For particles ofstatistically isotropic shape (on average, particles of the powderpossess no distinguishable orientation), such as spheres with roughsurfaces, the particle envelope volume is approximately equivalent tothe volume of cytosolic space required within a macrophage for completeparticle phagocytosis.

Swellable particles thus are capable of a longer term release of atherapeutic or other working agent. Following inhalation, swellablebiodegradable particles can deposit in the lungs (due to theirrelatively small size), and subsequently undergo swelling, slowdegradation and drug release, without the majority of the particlesbeing phagocytosed by alveolar macrophages. A drug can be deliveredrelatively slowly into the alveolar fluid, and at a controlled rate intothe blood stream, minimizing possible toxic responses of exposed cellsto an excessively high concentration of the drug. The swellableparticles thus are highly suitable for inhalation therapies,particularly in controlled release applications. The preferred massmedian aerodynamic diameter for swellable particles for inhalationtherapy is between 0.5 to 5 μm (prior to swelling). After swelling, theparticles have geometric sizes of greater than 6 μm volume mean diameterin the airways.

The particles may be fabricated with the appropriate material, surfaceroughness, diameter, density, and swelling properties for localizeddelivery to selected regions of the respiratory tract such as the deeplung or upper airways. For example, larger particles or more denseparticles may be used for upper airway delivery, or a mixture ofdifferent sized particles in a sample, provided with the same ordifferent therapeutic agent may be administered to target differentregions of the lung in one administration.

The swellable particles can be delivered by inhalation methods usingpropellant driven metered dose inhalers wherein hydrofluoroalkane and/oralkane liquefied gas propellants are used in these formulations withother excipients included for stabilization of the preparations. Drypowder inhalers can use swellable hydrogel particles prepared foraerosolization and inhalation. Use of a dry powder inhaler may requireblending with so called “carrier” particles (See Smyth and Hickey,Carriers in Drug Powder Delivery: Implications for Inhalation SystemDesign, American Journal of Drug Delivery, Volume 3, Number 2, 2005, pp.117-132). These carrier particles are typically lactose, sucrose,glucose or other particles that are blended with the swellable particlesfor aerosolization. Typically, the carrier particles are sized between50-500 μm as determined by sieve analysis and make up from 90-99% w/w ofthe powder placed in the inhaler for aerosol dispersion and inhalation.Dry powder insufflation, liquid spray systems, and nebulizers also canbe used to deliver the swellable particles.

Moreover, modulation of the aerodynamic particle size of the aerosolparticles can be used to target different regions of the airways.Attachment of targeting ligands on the surface of the swellableparticles can result in their localization at specific sites within therespiratory systems, such as for lung cancer a targeting ligand may beused to bind to a receptor that is overly expressed in that lung cancersuch as a folate receptor. Targeting also can be achieved by causing thehydrogel particle to change chemical bonding or conformation when theparticle is in a microenvironment that is unique to the disease site,such as in infection in the lung where inflammatory response of thelungs to the microorganisms causes higher concentrations of chemicalsand mediators that can cause the hydrogel to actively change its nature.This could be to cause pH sensitive changes in the hydrogel network sothat the drug loaded in the hydrogel particle is rapidly released whenpH decreases so as to concentrate the drug release to areas where thedisease is most pronounced.

Example 1 Synthesis of Peg Hydrogel

An eight-arm, hydroxyterminated PEGs with total number average molecularweights (Mw) of approximately 10 and 20 kD are acrylated to a degree offunctionalization exceeding 95% after azeotropic distillation followedby reaction with acryloyl chloride in the presence of triethylamine ashas previously been described by Elbert, D. L., et al., Self-selectiveReactions in the Design of Materials for Controlled Delivery ofproteins. Journal of Controlled Release, 2001. 76: p. 11-25. Hydrogelsare formed by mixing the chosen PEG-acrylate with either dithiothreitolor PEG-dithiol, Mw 3.1 kD at a 1:1 stoichiometric ratio of acrylates tothiols in 50 mM phosphate buffered solution (PBS, pH 7.8). Each reactantis dissolved separately in an aliquot of PBS. The amount of PBS isvaried to give the desired total precursor concentration (wt %) uponmixing. The two solutions are mixed vigorously in 1.5 mL plastic tubesand centrifuged to remove bubbles. The sealed tube containing themixture is placed at 37° C., and allowed to react overnight to ensurecomplete conversion.

There are significant toxicological and compatibility advantages ofusing such a reaction to form gels under physiological conditions fordrug delivery applications Peppas, N. A., et al., Physicochemicalfoundations and structural design of hydrogels in medicine and biology.Annu Rev. Biomed. Eng., 2000. 2: p. 9-29. The degradation rate of thepolymer is determined by various factors including the initial watercontent of the hydrogel network. This is initial water content iscontrolled by the crosslinker and the molecular weight of the PEGacrylate. Drug release from the hydrogel is determined by the relativesizes of the drug molecule and the mesh size of the crosslinked network.If drug size is assumed constant (though drug suspension particles couldconceivably be modulated in some cases), drug release can be modulatedby decreasing the mesh size. This is achieved by decreasing the lengthof the polymers (molecular weight).

Poly(acrylic acid-co-acrylamide) hydrogels were synthesized usingsimilar methods to those described by Chen, J. et al. “Synthesis andcharacterization of superporous hydrogel composites”, Journal ofControlled Release 65: pp. 73-82 (1999), the teachings of which areincorporated herein by reference.

Alternatively, the swelling particles for pulmonary drug delivery may beformed from polymers or blends of polymers with differentpolyester/amino acid backbones and grafted amino acid side chains, Forexample, poly(lactic acid-colysine-graft-alanine-lysine) (PLAL-Ala-Lys),or a blend of PLAL-Lys with poly(lactic acid-co-glycolicacid-block-ethylene oxide) (PLGA-PEG) (PLAL-Lys-PLGA-PEG) may be used.

In the synthesis, the graft copolymers may be tailored to optimizedifferent characteristics of the swelling particle including: i)interactions between the agent to be delivered and the copolymer toprovide stabilization of the agent and retention of activity upondelivery; ii) rate of polymer degradation and, thereby, rate of drugrelease profiles; iii) surface characteristics and targetingcapabilities via chemical modification; and iv) particle porosity.

Example 2 Dual Action Mucolytic-Therapeutic Drug Delivery Vector forCystic Fibrosis

From the moment an aerosolized drug is expelled from the metered-doseinhaler or nebulizer and enters the mouth, through its journey past thepharynx, down the trachea and bronchioles into the deeper recesses ofthe airways toward its site of action, and finally to its degradationand removal, aerosolized agents are under the influence of a multitudeof factors, which may be grouped into two general categories. Thosedeterminants which govern the deposition of the aerosolized agent ontothe airway lumen surface are termed physical properties, and include aparticle's diameter and density. The properties that determine the fateof the drug subsequent to its impaction on the luminal surface,including its absorption, metabolism and excretion, are referred topharmacokinetic factors.

Accordingly, the physical and pharmacokinetic factors of an aerosolizedparticle must be precisely tailored to complement one another as a meansof delivering the most effective dosage possible while simultaneouslyminimizing drug waste and circumventing undesired collateral reactions.

With these considerations in mind, this EXAMPLE pursuant to anotherillustrative embodiment provide a novel dual actionmucolytic-therapeutic hydrogel drug delivery vector for the treatment ofcystic fibrosis as a means of significantly improving the efficacy ofcurrent FDA approved cystic fibrosis therapeutics. Furthermore, althoughthis delivery system was initially designed specifically for cysticfibrosis, it can also serve as a therapeutic delivery vector for otherpulmonary disorders, including lung cancer, COPD, and asthma.

Due to the large amounts of pathogens and debris that we inhale witheach breath, the lung possesses multiple lines of defense to preventinfection and maintain homeostasis. From the trachea to the terminalbronchioles, an area collectively referred to as the central airways,the surface of the luminal epithelium is coated with a film of fluidthat is composed of a sol and gel phase and referred to respectively asthe periciliary and mucus layers. Each of these two layers play animportant role in keeping the lung clear of pathogens, and their precisecomposition is essential to the effective clearance of foreign particlesfrom the airways. The overlying thick and viscous mucus acts as abarrier to prevent the passage of inhaled pathogens and other foreigndebris to the underlying epithelial cells below. The periciliary fluid,while not a direct obstacle in the manner of the superjacent mucuslayer, is no less important to ensuring the lung is kept clean ofpathogens. Through the rhythmic and concerted beating of the epithelialcilia within the periciliary fluid, mucus is propelled in the cephalicdirection towards the pharynx and removed from the airway viaexpectoration or ingestion into the gastrointestinal tract, a processreferred to as the mucociliary escalator. Therefore, any aerosolizedagent remaining trapped within the mucus layer is carried along andsummarily removed from the respiratory tract.

The periciliary fluid is maintained at an optimal volume which ensuresthat only the tips of the cilia contact the overlying mucus, propellingit onward. The volume and ionic composition of the periciliary fluidlayer is the end result of the delicate balance between absorption andsecretion of H₂0 and ions (specifically Na⁺ and C^(I−)). This balance istightly regulated through the concerted action of ion channels locatedin the apical surface of pulmonary epithelial cells. Absorption of fluidis controlled primarily by the activity of the amiloride-sensitiveepithelial Na channel (ENAC). The absorption of osmotically active Na⁺ions from the periciliary fluid into the epithelial cells drives waterfrom apical to the basolateral surface of the cell, decreasing thevolume of the periciliary fluid. This action is countered by both theoutward rectifying chloride channel (ORCC) and the cystic fibrosistransmembrane conductance regulator (CFTR) channel, which allows thepassage of chloride ions out of the cell and into the lumen, carryingwater along and restoring the volume of the periciliary fluid.

In cystic fibrosis (CF), the most prevalent autosomal recessive geneticdisorder in Caucasians, the gene encoding the CFTR channel contains amutation which results in the translation of a defective protein. Themajority of the mutated CFTR protein is degraded in the endoplasmicreticulum by the 26S proteosome, and the small amount that does reachthe plasma membrane of the apical surface does not function properly andno longer allows the passage of chloride ions into the pulmonary lumen.This decreased transport of chloride ions results in a reduced amount ofwater entering the periciliary fluid, and since the Na⁺ channels arestill absorbing water at a normal rate, the volume of the periciliaryfluid is considerably reduced. The upper region of the cilia becomeembedded in the mucous layer, significantly hindering their movement andleading to the cessation of the mucociliary escalator. The mucus layerbecomes increasingly viscous and stagnant, allowing bacteria colonies torapidly accumulate throughout the lung. This increased infestationproduces a vigorous response from the immune cells of the body, whichover time significantly deteriorates the lung, and eventually results indeath.

The ideal solution would be to administer gene therapy that wouldreplace the defective copy in the chromosome with one that encodes for afunctioning CFTR channel. Unfortunately, while this is a very activearea of research, due to numerous setbacks gene therapy is still not aviable option, and treatment of the physical manifestations of thegenetic disorder remains the only route to treat CF. The current therapyconsists primarily of the mucolytic N-acetylcysteine, which severs thebonds between adjacent mucin glycoproteins and reduces the viscosity ofthe mucus, and administration of various antibiotics targeting thenumerous bacterial colonies infesting the lungs of CF patients. Thesetreatments are administered in aerosolized form and may be prescribedeither separate or in tandem. While this therapy does help alleviatesome of the complications of CF, it is an extremely wasteful procedureand much can be done to improve its efficacy while simultaneouslyreducing its cost in both time and money. For instance, the antibioticwill only be effective against the bacteria that it encounters as soonas it is deposited in the airways. Once the therapeutic contacts theviscous mucus it will be become entrapped, completely abrogating itsbactericidal activity. Conversely, no matter where the mucolytic landsit will be able to decrease the viscosity of the stagnant mucus in itsvicinity. However, the mucolytic alone does nothing to combat thenumerous bacterial colonies populating the airways, and which are thedirect source of the complications of CF.

This EXAMPLE provides swellable particles for delivering both themucolytic and therapeutic simultaneously within the same aerosol dose.In this manner the mucolytic will reduce the viscosity of the mucus andincrease the radius of diffusion of the therapeutic, bringing it intocontact with a greater number of bacteria and significantly improvingits effectiveness. The use of swellable particles to this end takes intoconsideration that any therapy administered this way, where thetherapeutic and mucolytic are delivered in their free (i.e.unencapsulated) form, will be short-lived. The immediate impact of thetherapy will permit the cilia to beat with increased frequency andimprove the function of the mucociliary escalator, which in turn willremove much of the mucus and bacteria, accompanied by the therapeutics,from the lung. However, as this treatment does nothing to address theunderlying genetic defect, the mucus will once again become stagnant andviscous, allowing the remaining bacteria an opportunity to divide andmultiply, so that soon after administration the condition in the lung isreturned to its pre-therapy state. On average, cystic fibrosis patientsmust spend approximately three hours per day self-medicating. Not onlydoes this strict and inflexible routine have a tremendous impact ontheir quality of life, it also gives rise to patient non-compliance withtheir prescribed therapy. Due to the rigorous nature of their treatment,a patient may either accidentally or intentionally forgo a day or two oftherapy, believing that such a brief absence of medication will not beseverely detrimental to their health. However, when fighting againstbacteria which require only hours to replicate this negligence may yieldsevere consequences. How then does one solve this paradox of alleviatingthe stringency of the therapy as a means of improving patient compliancewhile simultaneously ensuring the delivery of the prescribed dose ofmedication? The solution to this question is found in sustained releasetherapy, where the therapeutics are released continuously over a periodof days, if not weeks, allowing the patient to receive their prescribeddosage while increasing the interval between administration.

The use of swellable particles to this end takes into consideration theproblem that delivery via aerosolized particles completely ignores theaerodynamics which governs the deposition of aerosolized particles inthe lung. In aerosols a very important property is the aerodynamicdiameter (d_(a)) of a particle. The aerodynamic diameter is a methodused to standardize the aerodynamical properties of particles regardlessof shape, density, size, or actual diameter, and may be thought of asthe diameter of a water droplet having the same aerodynamic propertiesas described above. The most significant information derived from theaerodynamic diameter is that a particle with a given aerodynamicdiameter will be aerodynamically indistinguishable from other particlesof different size, shape, diameter or density having the sameaerodynamic diameter.

As the preceding paragraph demonstrates, the aerodynamic properties ofan aerosolized particle are heavily dependant upon both its' diameterand density, and the slightest alteration of either property will resultin the deposition of the particle in a different region of the lung.Therefore, attempting to deliver a mixture of mucolytic and therapeuticeither in their free form or encapsulated within discreet deliveryvectors will result in non-uniform deposition, since the two do not haveidentical diameters and densities and therefore will not possessidentical aerodynamic properties. Consequently, the EXAMPLE provides fordelivery of both the mucolytic agent and therapeutic agent togetherusing the swellable particles.

Moreover, the swellable particles provide a preferred CF therapeuticvector to provide sustained release of the therapeutic agent and alsocapability to carry both mucolytic and therapeutic agents. Furthermore,when designing a delivery vector for human consumption, it is essentialthat it is biocompatible, non-toxic, and non-immunogenic. The hydrogelparticles described above encapsulating the therapeutic and mucolyticagents within a crosslinked, hydrophilic polymer network achieves thesustained release of the agents. For example, when a hydrogel is placedin an aqueous environment it readily imbibes water and swells,stretching its polymer chains and producing numerous pores that willpermit the drugs within to diffuse out. Hydrogels are especiallyrelevant for drug delivery in the lung, as the initial surface that adelivery vector encounters is the viscous mucus, which is primarilycomposed of hydrophilic mucin glycoproteins, which readily attractwater. Therefore, hydrophilic polymer hydrogels are ideally suited tosiphon water away from the mucus, allowing the hydrogel to swell andrelease its therapeutic cargo.

The most important factor when developing a hydrogel network is theselection of a suitable hydrophilic polymer, and due its numerousbeneficial properties, polyethylene glycol (PEG) is described above.Among the attributes that make this FDA approved polymer attractiveinclude that it is non-toxic, biocompatible, non-immunogenic andstrongly hydrophilic, allowing it to draw water away from the mucus topromote hydrogel swelling. Furthermore, PEG exhibits almost no proteinadsorption, allowing it to elude the ubiquitous macrophages patrollingthe labyrinth of the pulmonary passages. These characteristics combineto make PEG the most promising candidate for employment in pulmonarydrug delivery.

The antibiotics used to treat CF include tobramycin and gentamycin,which are large, bulky, hydrophobic macromolecules. While theirhydrophobic nature prevents the antibiotics from being incorporated intheir free form, it permits them to be readily encapsulated within thecore of a liposome nanoparticle. The liposome nanoparticle in turnpossesses a hydrophilic surface, allowing it be easily incorporatedinside the hydrogel matrix. Furthermore, by adjusting the size of boththe liposome nanoparticle and the pores of the polymer matrix, therelease rate of the antibiotic from the hydrogel may be accuratelycontrolled.

The next step is to encapsulate the mucolytic (e.g. N-acetylcysteinedesignated NACS) within the hydrogel. Due to the small size of NACS andits hydrophilic nature, if it is simply loaded into the hydrogel in itsfree state, the moment the hydrogel begins to swell NACS will bereleased from the lung in an undesirable single, rapid burst. Toovercome this, the NACS is covalently bonded to the PEG polymer network,while still retaining its mucolytic activity. The advantage of thismethod is that as the hydrolytically labile ester bonds are broken NACSwill be released into the environment. This ensures that mucolytic willbe released throughout the entire lifetime of the hydrogel, providingthe sought after sustained release to accompany the diffusion of theantibiotic from the gel.

As previously mentioned, NACS is extremely soluble in water (100 mg/mL),thus precluding its incorporation into nanoparticles similar to thoseencapsulating the hydrophobic antibiotic.

The structure of N-acetylcysteine (pictured above) affords us twofunctional groups with which to create a hydrolytically labile bond toattach the mucolytic agent to the PEG-polymer network. The release rateof the therapeutic will depend on both the kinetics of thecleavage/degradation of the drug-network linkage (which is described byan appropriate rate constant), and upon the diffusion rate of the freemolecule from the matrix of the polymer network. The crosslinkeremployed in our hydrogels is dithiothreitol, which contains two thiolmoieties that readily react with our functionalized PEG octa-acrylateswithout the need of organic solvents at body temperature and biologicalpH.

However, since NACS contains only one thiol, it could either form adisulfide bond with another NACS molecule, or it may react with theacrylate group of the polymer. Either way, its function as a mucolyticagent would be significantly disrupted, if not abrogated entirely.

As shown in the reaction scheme (which is only showing the acrylate endgroup of the 8-arm PEG-acrylate coupling with the thiol functionalizedend of DTT), the initial step is the Michael addition of the thiol tothe acrylate group. Following this reaction, the next step is an esterhydrolysis in which an alcohol and a carboxylic acid are formed (theexact opposite of the Fisher esterification reaction). But as can beseen, the sulfur atoms (since what is formed is actually a dicarboxylicacid, since dithiothreitol contains two S—H groups per molecule) are nolonger in their reduced thiol forms. Instead, as shown below, they arebonded to two carbon atoms, forming C—S—C linkages, and no longer ableto participate in disulfide bonding with the thiol moieties of the mucinpolymers.

Therefore, control of the release of NACS can be achieved by reactingdithiothreitol with N-acetylcysteine, in perhaps a molar ratio of atleast 4:1 (DTT:NACS), prior to the crosslinking reaction with the 8-armPEG polymer. This will allow the thiols of DTT to react with the thiolof NACS. The advantage of NACS is that it is already an FDA approveddrug, and similar to other thiol compounds, will react readily withoutthe need for harmful solvents or reaction conditions. The structure ofthe cleaved molecule (following ester hydrolysis) is shown below:

As opposed to the previous compound, this molecule contains a disulfideS—S linkage, which readily competes with mucin polymers for theirdisulfide bonds, since the only exchange that is occurring is onedisulfide bond for another (requiring no significant cost in energy,unlike an exchange from a low-energy bond to a high-energy bond, whichwould be thermodynamically unfavorable and would possess a largetransition barrier) this exchange of disulfide bonds occurs under verymild conditions (essentially mixing the chemicals in a PBS buffer).

Although there will certainly be some NACS-NACS disulfide bonding, thisis not an irreversible linkage, and a greater amount of DTT will ensurethat the majority of the NACS will be bound to the crosslinker(theoretically, every molecule of NACS will yield one S—S activemucolytic bond). By increasing the amount of DTT used we can ensure thatthe hydrogel has the majority of the crosslinkable moieties occupied,thus maintaining structural stability, while simultaneously possessingenough DTT-NACS to exhibit a significant enhancement in hydrogelpermeability through the mucus. And while it is true that any DTT thatis bound to NACS will not be able to form a crosslink with another PEGmolecule, recent experiments show that firm hydrogels can be formed withat least 30% concentration of crosslinker compared to the polymer (thatis to say that there are enough DTT molecules to theoretically occupy30% of the PEG acrylate groups).

Synthesis of Biocompatible Hydrogels from 8-Arm PEG Acrylate ContainingRhodamine and N-Acetylcysteine

The initial synthesis of the acrylated 8-arm PEG (Mw=10 kDa and 20 kDa)was based upon the work of Hubbell et al. Journal of Controlled Release76:11-25 (2001) as described above. For example, ten (10) g of 8-arm PEG(20 kDa, Nektar) was dissolved in 200 mL of toluene and distilledazeotropically for 2 hrs. The resulting solution was then allowed tocool to 50 C under argon. Two (2) mL of triethylamine was added to 50 mLof dichloromethane which was then added to the reaction solution. Anamount (1.3 mL) of acryloyl chloride was then added dropwise and thereaction proceeded under argon in the dark for 20 hrs. The resultingopaque pale yellow solution was then filtered multiple times untilclear. Anhydrous sodium carbonate was added to the solution and stirredfor two hours to remove any water that was present. The solution wasfiltered to remove the sodium carbonate and was then evaporated underreduced pressure. Diethylether was then added to the solution and thereaction flask was placed in an ice bath to allow the product (acrylatedPEG) to precipitate. The product was collected by filtration andrepeatedly washed with diethylether. The average yield was around 85%.

Firm and stable PEG hydrogels can be formed with an acrylate:thiolstoichiometric ratio >1 and with a thiol amount as low as 60% (0.60ratio of thiol/acrylate) of the amount of acrylate moieties present.Generally, dithiothreitol, having two thiols per molecule, formscrosslinks between polymer molecules to form the hydrogel network. Toform hydrogels with N-acetylcysteine mucolytic agent covalently bound tothe hydrogel network, we reacted 5 mg of dithiothreitol with 1.1 mg ofN-acetylycysteine, each dissolved in 20 ul of 1×PBS (7.4 pH).N-acetylcysteine has the capacity to form a bond between it's thiolgroup and either the acrylated polymer or to dithiothreitol via S—Sdisulfide bonds formed from two thiol groups. Stiochiometrically, at theabove mentioned concentrations of reactants, if all of the thiol groupson the N-acetylcysteine are each bound to one unique dithiothreitolmolecule, there will still remain enough free thiol groups ondithiothreitol molecules to form sufficient crosslinking between thepolymermolecules such that stable hydrogels are formed. This solutionwas added to 0.160 g of acrylated 8-arm PEG (20 kDa) dissolved in 200 μlof 1×PBS (pH 7.4). To this solution was added 10 μl of Rhodaminetherapeutic agent (40 mg/mL 1×PBS (pH 7.4)). Seventy (70) μl aliquots ofthe solution was placed between microscope slides coated with SigmaCote(Sigma Chemical Co.), and separated by 1 mm spacers. The gels wereallowed to cure for 24 hours in a humid environment at 37 C. The curedgels were milled after drying, in a micro-ball mill (from Dentsply Rinn,Elgin, Ill.) cooled using liquid nitrogen, to produce swellableparticles having volume mean diameters of between 1.1 and 3 μm and aspan of 2.2 (span=(D90−D10)/D50) where D50 is median diameter and D10and D90 are respective 10^(th) and 90^(th) percentile diameters (e.g.for D10, 10% of particles are less than this diameter).

As discussed above, in this EXAMPLE, dual delivery of both mucolytic andtherapeutic agents would serve to significantly enhance theeffectiveness of the present CF therapy by increasing the radius ofdiffusion of the therapeutic, allowing it to contact a larger number ofbacteria, while simultaneously improving the function of the mucociliaryescalator. However, this dual-action aerosolized hydrogel is not limitedfor use to cystic fibrosis therapy. As previously mentioned, duringphysiological conditions the lumen of the pulmonary passages are coatedwith a layer of mucus serving as a barrier to inhaled particles. Thismucus is continuously removed via the mucociliary escalator, requiringapproximately 10 hours to eject inhaled debris from the furthest reachesof the lung, and much less time to clear the upper passages where themajority of malignant tumors dwell. Incapable of distinguishing betweentherapeutic aerosols and pathogens, the mucociliary escalator expelsfriend and foe alike with equal vigor. Accordingly, any inhaledtherapeutic has only a brief window in which to penetrate the mucousbarrier, attain the underlying epithelium, and deliver its medicinalcargo, else it forever loses any opportunity for efficacy. Furthermore,when one also considers that the tumor is not uniformly distributedthroughout the lung and that there is only a specific region where thetherapeutic will be effective, the aforementioned brief window of actionis further narrowed. By controlling the aerodynamic properties of aparticle via its diameter and density, an aerosol can be tailored sothat the majority of the particles will arrive at the desired locationin the lung. However, once the particle lands on the surface of thelumen, the onus is entirely upon the particle to penetrate the mucusprior to its expulsion from the lung.

Therefore, it becomes evident that the quicker a particle can passthrough the viscous mucus, the greater its chances to provide abeneficial effect. By combining a mucolytic with a cytotoxic agent,corticosteroid or bronchodilator for the treatment of lung cancer, COPD,and asthma respectively, the efficacy of the treatments will beincreased and the amount of drug wasted via the action of themucociliary escalator will be markedly reduced.

Targeting molecules can be attached to the swelling particles viareactive functional groups on the particles. For example, targetingmolecules can be attached to the amino acid groups of functionalizedpolyester graft copolymer particles, such as PLAL-Lys particles.Targeting molecules permit binding interaction of the particle withspecific receptor sites, such as those within the lungs. The particlescan be targeted by attachment of ligands which specifically ornon-specifically bind to particular targets. Exemplary targetingmolecules include antibodies and fragments thereof including thevariable regions, lectins, and hormones or other organic moleculescapable of specific binding for example to receptors on the surfaces ofthe target cells.

Although the invention has been described above with respect to certainillustrative embodiments, those skilled in the art will appreciate thatchanges, modifications and the like can be made thereto withoutdeparting from the spirit and scope of the invention as defined in theappended claims.

1. Swellable particles for delivering a working agent to the pulmonarysystem, the particles comprising: a plurality of biodegradable particleseach formed from a polymer network, each of the plurality ofbiodegradable particles having a mass median aerodynamic diameter notexceeding 5 μm, the particles being swellable by hydration to a sizethat is greater than 6 μm volume mean diameter; and a working agententrapped in the polymer network of each of the plurality ofbiodegradable particles. 2-3. (canceled)
 4. The particles of claim 1,wherein the working agent is entrapped in nanoparticles, wherein thenanoparticles are incorporated in the biodegradable particles.
 5. Theparticles of claim 1, wherein the working agent is chemically bondedwith the polymer network of the biodegradable particle.
 6. The particlesof claim 1, wherein the working agent comprises one or more of atherapeutic treating agent, a diagnostic agent, a prophylactic agent, oran imaging agent.
 7. The particles of claim 6, wherein the working agentcomprises a mucolytic agent.
 8. The particles of claim 1, wherein theworking agent comprises at least one of an antibiotic agent, a cytotoxicagent, an RNA interfering agent, and a gene. 9-11. (canceled)
 12. Theparticles of claim 1, wherein the working agent comprises multiplecytotoxic agents.
 13. The particles of claim 1, wherein the workingagent comprises a cytotoxic agent and an RNA interfering agent.
 14. Theparticles of claim 1, wherein at least 90% of the particles have anaerodynamic diameter of 5 μm or less and swell to a size greater than 6μm volume mean diameter.
 15. The particles of claim 1, wherein thepolymer network of the biodegradable particles comprises a materialselected from the group consisting of a biodegradable natural polymer, asynthetic polymer, a protein, and a carbohydrate, or combinationsthereof.
 16. The particles of claim 1, wherein the plurality ofbiodegradable particles comprise hydrogel particles.
 17. The particlesof claim 1, including a coating on the plurality of biodegradableparticles that controls a rate of particle swelling.
 18. The particlesof claim 17, wherein the coating comprises an excipient selected fromthe group consisting of a carbohydrate, a lipid, a protein, or abiocompatible salt of sodium, potassium, calcium, magnesium or lithium.19. The particles of claim 17, wherein the coating comprises a targetingagent for binding to receptors or to a target within a diseased site.20. An aerosol comprising the swellable particles of claim
 1. 21. Theparticles of claim 1, wherein the polymer network of each of theplurality of biodegradable particles comprises a cross-linked polymernetwork.
 22. The particles of claim 21, wherein the cross-linked polymernetwork physically entraps the working agent in the polymer network. 23.The particles of claim 21, further comprising a crosslinker to formcrosslinking of polymers within the polymer network.
 24. The particlesof claim 23, wherein the crosslinker comprises dithiothreitol,